When a sample or a human body or its part (further generally mentioned as “a sample”) is subjected to uniform polarizing magnetic field (usually high-strength stationary magnetic field B0=1 . . . 3 T, up to 7T, used for medical imaging, or low-strength—B0=0.05 . . . 0.5 T, used for security imaging as well as for nuclear magnetic resonance characterization of samples), the individual magnetic moments of the spins in the subject tissue attempt to align with this polarizing field creating an equilibrium magnetization M0 in the tissue, which is proportional at equilibrium to the magnetic field. While the stationary magnetic field is applied along z-axis of MRI apparatus, the tissue is subjected to cyclically applied excitation field B1 in the x-y-plane generated by a resonance frequency (RF) coil, particularly to at least one electromagnetic near-resonant radio-frequency pulse (RF excitation pulse applied with resonant Larmor precessing frequency, which is proportional to B0, 42.58 MHz/T for hydrogen). With RF excitation pulses applied in the presence of a slice-selection gradient, the net aligned (longitudinal) magnetization Mz may be tipped into the x-y-plane producing transverse magnetization Mxy in a selected 2D slice. The longitudinal magnetization Mz tends to M0 with characteristic spin-grid relaxation time T1. The transverse magnetization is susceptible to Landau-Lifshitz-Gilbert damping; Mxy relaxation in the tissue is characterized by spin-spin relaxation time T2. The energy content of the RF pulse determines the amount of excited spin which is capable of emitting a MR signal. The transverse magnetization is characterized by its ratio to the longitudinal magnetization, or by a flip angle α of a RF excitation pulse. The MR signal emitted by the excited spins after so-called echo time provides information related to the properties of the tissue; the MR signal may be acquired to characterize the sample forming data sets from its different locations and to form, in such a way, an image, which may be, for example, a T2-weighted image. Typically, the region to be imaged is scanned by a sequence of varying magnetic field gradients, which are spatially encoded in all three spatial directions using gradient coils with necessary gradient waveforms applied. The resulting set of received echo signals acquired by appropriate RF reception coils in each measurement cycle is digitized and processed using one of well-known image reconstruction techniques.
A tissue layer (slice) location is selected in a spatial direction defined as a slice-selection direction (S-direction usually corresponding to z-direction) by applying in this direction a layer-selection (slice-selection) Gx=GS gradient simultaneously with the excitation RF pulse. For the purpose of achieving local resolution in a second spatial direction defined as a phase-encoding direction (P-direction usually corresponding to x-direction), typically a so-called phase-encode Gx=GP gradient is temporary imposed in this direction, which ensures dephasing after the excitation RF pulse and rephasing prior to the detection of the MR signal of oscillating spins along said direction. The phase-encoding gradient is turned on and off after each excitation RF pulse with different gradient magnitude to encode/decode the received signals in the x-direction before data is collected with the readout gradient applied. In a third spatial direction defined as a readout direction (R-direction usually corresponding to y-direction), typically a so-called frequency-encode Gy=GR gradient is imposed, usually synchronized with the phase decoding and with signal acquiring. To encode/decode signals in the y-direction, the signals are detected in the presence of so-called frequency-encode or readout gradient, to enable collecting a line of data points with different precession frequency corresponded to a single RF excitation pulse and their mapping into a spatial frequency domain known as k-space. Since phase and frequency are separately dependent upon a location along P- and R-direction correspondingly, it is possible to reconstruct a MR image in xy-plane sampling k-space signal prior to Fourier transformation.
The T2-weighting sequence applied to the same slice is repeated with different phase encoding with period TR, a repetition time. TR determines the spin echo sequence duration; the longer the TR (up to and over 2000 milliseconds), the more complete the longitudinal magnetization regrowth and correspondingly the greater initially excited transverse magnetization Mxy. Associated with greater Mxy and longer echo-time TE (80 to 140 ms), the different tissues are better highlighted according to their spin-spin relaxation time T2 (which should exceed TE and may be as low as 150-200 ms). However, the long repetition time results in long acquisition time.
The diffusion-weighted MR imaging (DW-MRI or DWI) is a T2-weighted imaging method, which may be used for early detection of brain infarcts as well as for detection of liquid explosives and for characterization of hydrocarbon-bearing formation samples. The at least one pair of diffusion-weighting gradient pulses are applied along one axis or consequently along plural gradient axes, with time interval between beginnings of first and second gradient pulses called diffusion time interval A. The diffusion-weighted preparatory sequences ensure that a molecular movement which occurs in the observed voxel as a result of diffusion processes produces an identifiable attenuation of the T2-weighted MR signal, according to the following equation known as a partial solution of Bloch-Torrey differential equation for transverse magnetization Mxy:S=S0exp(−bD),  Eq. (1)where, S, S0 are signal intensities with and without applying diffusion gradients; D is apparent diffusion coefficient (ADC); b is a variable dependent upon parameters of the respective preparatory sequence and the gyro-magnetic ratio for protons γ=2.67·108 rad/(s·T); in the most common case of one pair of diffusion-weighting gradient pulses of magnitude G and duration δ, it's determined by the formula:b=(δ7G)2(Δ−δ/3).  Eq. (2)
In order to obtain effective diffusion-weighting, i.e. effective identifiable attenuation determined by diffusion, and to determine ADC with acceptable accuracy, generally b-value in the range of 500-2000 s/mm2 is required; diffusion times in the state-of-the-art techniques at restricted gradient values are usually of order 20-40 ms. Such diffusion time order corresponds to diffusion path (Einstein's mean radial displacement) <z>=(δDΔ)1/2 of order 10 μm, which determines boundary restrictions (blur).
The standard diffusion-weighting method has its origins in the Stejskal-Tanner experiment (see: “Spin-diffusion measurements: spin echoes in the presence of a time-dependent field gradient” by E. O. Stejskal, J. E. Tanner, J. Chem. Phys., Vol. 42, 1965, p. 288-292). The spin-echo sequence comprises a 90° excitation and 180° refocusing RF pulses and at least one pair of diffusion-weighting gradient pulses. As a method of refocusing the MR signal for its detection, the so-called spin echo-signal is produced by refocusing the RF magnetic field by means of an additional refocusing RF pulse, commonly with flip-angle 180° (or, for example, two 90° RF refocusing pulses in a case of so-called stimulated echoes), which is excited in the presence of the same slice-selection gradient, particularly, in a half spin-echo time TE/2 after the initial excitation RF pulse, where another half spin-echo2 time TE/2 is a time between the refocusing RF pulse and MR signal detection. Spin-echo sequence is well combined with Echo-Planar Imaging (SE-EPI) method of image acquisition comprising acquiring multiple k-space lines in a single shot (rapid sequence of echoes with alternation of the polarity of the read gradient during 20-80 ms) that avoids decreasing SNR (signal-to-noise ratio). The signal intensity S is determined at DWI for every voxel accordingly to Eqs.(1), (2) by choice of gradients and diffusion times within echo time TE during which the peak signal is obtained. For the sample characterization, the scan results may be compared at different diffusion-weighting characteristics. For obtaining comparable results of different DWI experiments, the motion compensation is desirable.
However, for modern high-resolution diffusion sequences with standard SE-EPI protocol characterized by long readouts, total acquisition time approaches to 14-20 min and more, which is clinically unacceptable; the higher image resolution comes at a cost of sufficiently increased scan time and the blur associated with T2* spin-spin decay. The image spatial resolution is low, as the number of consecutively detectable echoes at given echo time TE is very limited by the naturally short effective spin-spin relaxation time T2*, which relaxation includes irreversible decay caused by static field magnetic susceptibility inhomogeneities. From the other hand, the gradient values are limited by the gradient coil performance limits and the safety issues; at that achieving sufficient b-value requires increasing Δ and TE. At larger Δ, the image becomes more sensitive to the higher-order derivatives of the body bulk motion; therefore, the motion compensation is desirable, especially at steep diffusion gradients.
In the alternative gradient-echo sequence without the refocusing RF pulse, the pair of gradient pulses has a form of two lobes with opposite polarity (bipolar gradient—BG pulses). The polarity inversion of the magnetic field gradient is used for refocusing after dephasing created by the first polarity part of the BG pulse. Due to large phase shifts from small patient bulk translations, the gradient-echo DWI is exquisitely sensitive to motion because of absence of polarity and phase velocity reversing by the 180° refocusing RF pulse. For example, at BG diffusion-weighting pulse characteristics G=40 mT/m, δ=20 ms, Δ=30 ms, value b=580 s/mm2 is achieved. At that, in the event of macroscopic movement x of the human body, π-order phase shift in the MR signal is occurred at the local velocity of order 0.3 mm/s, which shift is generally determined by the formula:φ˜γGδ.  Eq. (3)
As no 180° rephasing pulse is used, the relaxation due to fixed causes is not reversed; the loss of signal resulting from low effective spin-spin relaxation time T2* result in lower signal-to-noise ratio SNR. For decreasing boundary blur and motion artifact effect, the diffusion time interval A may be limited, then obtaining quality images may be complicated. Ultimately, the motion compensation should be used.
An example of the steady-state gradient-echo refocusing method is given by the Diffusion-Weighted Steady-State Free Precession (DW-SSFP) imaging method (see: “Steady-state diffusion-weighted imaging: theory, acquisition and analysis” by J. A. McNab and K. L. Miller, NMR Biomed., Vol. 23, 2010, p. 781-793). Steady-state imaging involves the repeated application of identical sequences of low angle RF excitation pulses and a DWI gradient waveforms synchronized accordingly to net magnetization rotation; at that each sequence is unbalanced with respect to the first derivative of bulk motion. At lower flip angle α, the recovery of the residual longitudinal magnetization for a given T1 will be sufficiently complete at shorter TR; at that shorter echo time TE is required to provide sufficient signal strength at acquisition.
The longitudinal magnetization gradient created by DWI gradient waveform coupled with low-angle RF excitation (usually α=15-30°) determines the drift-diffusion effect multiplying the magnetization mobility (diffusivity) in the longitudinal direction, which can be easily detected as a greater attenuation of T2*-weighted signal. (Strictly speaking, the quasi-steady-state free precession with cyclical drift of the spins takes place in a case of DW-SSFP with longitudinal diffusion gradients; however, consistently with established terminology, below it's referred as a steady state with cyclical drift or free precession with cyclical drift). While the complicated interdependence of the diffusion terms and the nondiffusion terms (especially in the case of unbalanced gradients) make it difficult to build intuition for how DW-SSFP pulse sequence will behave for a given parameter choice and/or tissue type and to find required diffusion-weighting parameters, experiments show an apparent diffusion coefficient anisotropy at steady state—increasing ADC in z-direction as high as up to 2-10 times, which may be explained by drift-diffusion effect in the presence of the magnetic field gradient co-directed with polarizing magnetic field and longitudinal magnetization.
However, DW-SSFP is practically very difficult to apply to in vivo imaging due to strong sensitivity to motion at repeated excitation technique. Because steady-state sequences conserve residual transverse magnetization, these sequences are extremely sensitive to motion artifacts arising due to a phase shift caused by translation in the diffusion gradient direction and to a phase gradient caused by the tissue deformation in the diffusion gradient direction, or by rotation in the plane perpendicular to the diffusion gradient direction.
Multi-slice MRI imaging is known from the prior art. In the method of multi-slice SSFP facilitating multi-slice coherent imaging according to U.S. Pat. No. 7,514,927 to Herzka and Winkelmann, issued 2009 Apr. 7, for a plurality of mutually perpendicular steady spin-state slices, time-domain multiplexed readouts are performed with applying phase-encode gradients in the mutually perpendicular directions; acquired magnetic resonance data in each readout correspond to one of the steady spin-state slices. Radio frequency excitation pulses are interleaved with the time domain multiplexed readout pulses to maintain the steady spin-states of the slices. The acquired magnetic resonance data are reconstructed to produce reconstructed images corresponding to the plurality of slices. In such a way, the multi-slice coherent imaging reduces general acquisition time.
However, the mutually-perpendicular multi-slice imaging is related to the noise amplification; it's extremely sensitive to motion artifacts, especially at diffusion-weighting; its applicability for DW-SSFP is strictly limited.
In the method for simultaneous multi-slice MRI according to U.S. patent application Ser. No. 12/761,314 to Setsompop and Wald, published 2011 Oct. 20 (2011/0254548 A1), image data is acquired simultaneously from multiple slice locations using a RF coil array. Following RF excitation, at least one readout gradient is applied along frequency-encode direction in order to form echo-signals. For separating of aliased pixels at multi-slice acquisition, a modified SE-EPI sequence includes a series of slice-encode gradient “blips” contemporaneously with phase-encoding blips in the presence of alternating readout gradients common to EPI sequences. The slice-encoding blips are designed to impart a phase shift accruals to the formed echo-signals; the image data corresponding to each sequentially adjacent slice location is cumulatively shifted by a fraction of the imaging FOV in the phase-encoding direction (field-of-view, which for a conventional brain image may be equal to ˜24 cm). This fraction FOV shift in the image domain provides separation of the aliased signal data with substantial suppression of pixel tilt and blurring at the image reconstruction. For diffusion-weighting in a selected direction, two diffusion gradients are applied prior to applying the at least one readout gradient.
However, this method based on a phase shift by a fraction of FOV is very sensitive to a phase shift due to the second derivative of patient bulk motion; the motion compensation is required for avoiding motion artifacts.
A useful gradient echo may be generated using Fast Low-Angle Shot (FLASH) sequence comprising crusher gradient pulse for eliminating in each cycle residual transverse magnetization, which, however, entails the high noise level. The higher SNR and CNR are provided by the Fast Imaging with Steady Precision (True-FISP) technique that uses fully balanced gradient waveform. The elements of both techniques are used in method of reducing imaging scan times according to U.S. Pat. No. 8,076,936 to Borthakur e.a., issued 2011 Dec. 13, the sequence comprises a T1p preparation period including applying a first non-selective 90x° magnetization-flipping RF pulse, one or more phase-alternating spin-locking 90y° RF pulses to provide T1p contrast reducing image artifacts resulting from B1 inhomogeneity, a second non-selective 90−x° RF magnetization-flipping pulse to return the magnetization to the longitudinal axis (with characteristic T1p, relaxation time); followed by a balanced gradient echo sequence for image acquisition including applying at least one α/2−x excitation pulse followed by balanced pairs of phase-alternating spin-locking αx and α−x pulses and magnetization-spoiling crusher gradient pulse to reduce blurring caused by the approach to the steady-state; followed by acquiring multiple lines of k-space and allowing magnetization to relax toward equilibrium magnetization (with possible relaxation period shortening to the time 0.3-0.4 s instead of 2-4 s and using model of magnetization saturation).
However, applying the balanced pairs of phase-alternating spin-locking impulses and additional crusher gradient pulse is associated with higher noise and significant signal loss incurred when the balance is not fully restored; the motion compensation is desirable.
In vivo intra-voxel incoherent motion MRI technique may be used, particularly, for the blood flow measurement; a perfusion tensor may be obtained. The method of perfusion imaging according to U.S. Pat. No. 7,310,548 to Van Den Brink, issued 2007 Dec. 18 comprises performing a first magnetic resonance data acquisition at a zero or very low sensitivity (b1=0) value, performing a second set of at least six magnetic resonance data acquisition in different directions at low sensitivity (b2<50 s/mm2) values; determining the perfusion tensor based on the magnetic resonance data acquisitions. For improving contrast, the method embodiment includes performing the second and a third sets at higher sensitivity (b2=600-1200 s/mm2, b3=100-400 s/mm2); determining diffusion coefficient tensor to provide a diffusion signal component; eliminating the diffusion signal component to provide perfusion signal component. The perfusion tensor computation based on the decay measuring at different sensitivities is determined by bi-exponential MR signal decay function:S/S0=fexp(−bP)+(1−f)exp(−bD),  Eq. (4)
where, P is a perfusion constant (typically P=0.05-0.08 mm2/s); D—diffusion constant (typically D=0.002 mm2/s); f is a fraction of fluid (flowing material, representative fraction of blood content in a voxel).
Low-sensitivity spin-echo diffusion-weighted sequence with single signal acquisition allows distinguishing perfusion characterized by higher-derivative flow patterns; at that the data of computed image maps of apparent diffusion coefficient (ADC) may be fitted to a stretched model of a tissue heterogeneity index mapping. The perfusion-imaging apparatus according to U.S. Pat. No. 7,310,548 comprises MRI data acquisition device, a display unit, and a computer system programmed to perform MRI data acquisitions with at least two different low b-values and with at least one higher b-value, and is further programmed to determine diffusion signal component of the low b-value acquisitions based on the higher b-value acquisition.
However, the blood flow rates vary in a broad diapason and can't be estimated directly based on this sequences; it's time-consuming to select the sensitivity values in a working range eliminating image flow artifacts while providing resulting signal attenuation sufficient for diffusion-weighting analysis. While useful signal not containing the flow velocity derivatives is analyzed, practically it's difficult to distinguish perfusion flow from higher-order bulk motion derivatives and to eliminate diffusion signal component. For obtaining comparable results in different diffusion-weighting experiments, the motion compensation is desirable.
The first acquisition data set with no or very low diffusion-weighting may be used as a target image, particularly, to construct a phase map in k-space; at that spatial misalignment may take place. The phases from the higher weighted data may be combined with the phase map of the first set with eliminating a motion induced phase shift and forming the adjusted data set using correlation coefficient similarity measure (see: “Correction of motional artifacts in diffusion-weighted images using a reference phase map” by A. M. Ulug et al., Magnetic Resonance in Medicine, Vol. 34, No. 2, 1995, pp. 476-480).
However, the motion artifacts are very difficult to avoid at the phase shifts more than π(180°), which is determined by diffusion sensitivity accordingly to Eq. (3); said phase shift is decoded as a space parameter.
In another method of correcting spatial misalignment using normalized mutual information similarity measure, the images are registered to a normalized template using a mutual information-based registration technique and employing a spatial transformation model containing state-of-the-art dependencies to correct the eddy current-induced image distortion and rigid body motion in three dimensions (see: “Comprehensive Approach for Correction of Motion and Distortion in Diffusion-Weighted MRI” by G. K. Rohde et al., Magnetic Resonance in Medicine, Vol. 51, No. 1, 2004, pp. 103-114).
However, sufficient patient motion and tissue deformation may lead to the misregistration of subsequent image nodes (phase-encode lines in 2D images or slices in 3D images).
In the method of acquiring MRI image according to U.S. Pat. No. 6,842,000 to Norris and Driesel, issued 2005 Jan. 11, double spin-echo T2-weighted preparatory sequence comprises a 90x° RF excitation pulse followed by two pairs of bipolar diffusion gradients (BG pulses) and a 180y° RF refocusing pulse between two gradients of each pair, at that each of the 90x° excitation pulse and the 180y° refocusing pulses are applied in the presence of the layer-selection Gz=GS gradient. The double spin-echo occurs and two gradient navigator echoes are generated consecutively during the second spin echo of the preparatory sequence; the navigator echoes contain bulk motion characteristics measured as an interference variable. The disturbing phase changes are measured at points of corresponding spin echo maximums and then a navigation signal is analyzed and the phase characteristics of transversal magnetization are corrected by compensation pulses. At that a homogeneous magnetic field pulse dimensioned to compensate the zero-order phase shift and magnetic field gradient pulses dimensioned to compensate the phase gradients measured by the navigator signal interference are applied in the direction of the stationary longitudinal magnetic field and in the correspondent spatial directions. In a case the direction of the diffusion gradients is the same as S-direction, P- and R-components of the phase gradient are encoded separately in the navigator echoes; the corrective gradient pulses are applied in the P- and R-direction respectively. Because spatial encoding-decoding is not done during T2-weighted preparatory imaging sequence, a 90−x° driven equilibrium Fourier transform pulse is applied at the time of the last spin-echo rotating the transversal magnetization back onto the longitudinal axis, its phase may be shifted to compensate the measured motion-related phase shift. It may be followed by an image acquiring sequence with N consecutive periodic cycles with phase encoding-decoding P- and R-gradient pulses changed from cycle to cycle to fill the various lines of the k-space used to reconstruct the image, for example, the FLASH sequence of partial experiments with a short repetition time TR<<T1 in each time containing an excitation with the flip angle α<90°.
Apparature for the MRI data acquisition according to U.S. Pat. No. 6,842,000 comprises a magnet for generating constant B0 field, an additional coil for generating parallel homogeneous magnetic field Bz, gradient coil sets for generating gradients Gx, Gy, Gz, a power supply unit comprising a RF generator and gradient coil power supplies, a RF coil and a selection unit for selecting the MRI signals, a control unit comprising a correction control path. The control unit controls selection of MRI signals and uses the interference variables of the navigation signal for the correction of bulk motion artifacts by changing the characteristics of the transversal magnetization; the correction control path configuration allows upon activation thereof to apply a homogeneous magnetic pulse in the direction of the stationary magnetic field dimensioned to compensate for the phase shift and to apply magnetic gradient pulses to compensate the phase gradient for the transversal magnetization.
In the navigator echo technique, displacement of the moving subject is measured for each location by the shift of the image space navigator echo compared to a reference echo. However, it's difficult to correctly assign phase shift to the correspondent location because the motion-disturbed phase characteristics of the navigator echoes are incompatible by phase in k-space. The data acquisition by followed FLASH sequence is more time-consuming than widely used SE-EPI method; at that the image spatial resolution is impaired.
Summarizing, using the known high-productivity MRI techniques, such as SE-EPI, may lead to motion-related phase misregistration and artifacts. Particularly, the known diffusion-weighting MRI methods achieve required quality at the cost of high diffusion-weighting gradient sensitivity, long diffusion and repetition time, generally resulting, for gradient-echo sequences, in high sensitivity to bulk motion and vibration and, for spin-echo sequences, to higher-order motion derivatives. The known methods of motion compensation don't sufficiently achieve the goals of artifact elimination at sufficient bulk motion, of distinguishing bulk motion from perfusion, etc.